Recent advances in the field of medical imaging have greatly facilitated the transition from technologies used to accurately image structures inside the human body to technologies sensitive enough to provide functional and biological information at the cellular and the molecular level. PET is considered to be one of the most sensitive in-vivo molecular imaging modalities despite its significantly inferior spatial resolution compared to imaging modalities such as computed tomography (CT) and magnetic resonance imaging (MRI). The improvement of PET detector technology is an active field of research and efforts are focused on addressing the limits in spatial resolution and sensitivity achieved in PET.
FIG. 1 shows a schematic block diagram of a PET imaging system in which the present invention can be implemented. In PET imaging, positrons are emitted from a radiopharmaceutically doped organ or tissue mass of interest of a patient 100. The positrons combine with electrons and are annihilated and, in general, two gamma photons which travel in diametrically opposite directions are generated upon that annihilation. A PET acquisition is based on the coincident detection of many pairs of simultaneous anti-parallel photons following the annihilation of the positron. The detection is performed by a detector 110 which comprises a plurality of detector element pairs which are placed around the imaged object, typically in a ring geometry. Opposing crystal detectors, which each scintillate upon being struck by a gamma photon, are used to detect the emitted gamma photons. Coincidence detection may be performed by a coincidence processing unit 120. By identifying the location of each of two essentially simultaneous gamma interactions as evidenced by two essentially simultaneous scintillation events, a line in space along which the two gamma photons have traveled (a “line of response,” or “LOR”) can be determined. The LORs associated with many million gamma interactions with the detectors are calculated and reconstructed to generate an image of the organ or tissue mass of interest in an image reconstruction unit 130. Thus, in PET imaging systems the detection of incident radiation is achieved by a two-step conversion of the annihilation photon energy to visible light in a scintillation material and to electric charge in the detector 110.
Time-of-flight PET (ToF-PET) is an advance over traditional PET that exploits the arrival time difference in detection of two photon events and correlates it to the position of the annihilation point with respect to the center of the field of view (FoV). The flight time difference between the two detected photon events is in a first approximation related to the object position along the line connecting the two detector elements: dx=c/2dt. The benefit in image quality depends on the time-stamp jitter. The effect is fairly dramatic, as each time the jitter is reduced by a factor of two, the patient acquisition time can also be reduced by a factor of two.
There are several types of photo detectors available: The first family of photo detectors covers the vacuum tubes: Photomultiplier tubes (PMTs) with a fairly large detector area of several square centimeters, and multi-anode PMTs which provide position information of a few millimeters, allowing pixilation in the millimeter range.
The photomultiplier tube (PMT) is a photo-detector type commonly used for scintillator readout in numerous applications including medical imaging. In ToF-PET, PMTs are often used for sub-nanosecond time resolution. The basic component of a PMT is a vacuum tube consisting of a photo-cathode, several electrodes called dynodes and an anode. The photocathode is a photo-sensitive electrode that emits charge (electrons) for a number of incident optical photons absorbed with a given quantum efficiency (QE). Between the cathode and the anode a bias voltage in the kV range is applied to facilitate the generated electron transport and amplification from the cathode to the anode. Under the influence of a high potential the generated electrons from the photocathode drift and successively encounter the dynode stages. At every dynode stage the incident electrons have gained sufficient energy to create secondary electron emissions from collisions with the dynode, thus resulting in a large electron cloud at the anode.
The second family of photo detectors is silicon based and incorporates either avalanche photo-diodes (APDs), analog or digital silicon photomultipliers (SiPMs) which are based on multi element avalanche photodiodes driven in Geiger Mode. All silicon detectors allow designs of PET/SPECT detectors with a small pixilation in the millimeter range. But only the analog and digital SiPMs (dSiPMs) allow sub nano-second timestamping to exploit the benefits of ToF-PET.
Furthermore, silicon photomultipliers (SiPMs) have been introduced to address shortcomings of PMTs to realize smaller pixilation. A novel technological advance in the field of semiconductor photo-detectors has been recently developed and involves the integration within the SiPM sensitive area of basic processing electronics thus reducing the need for external processing electronics. Each micro-cell of the array is connected to an integrated counter (for extraction of energy information) and an integrated time to digital converter (TDC) for extraction of time information. This alternative SiPM design is known as “digital SiPM” or “dSiPM” and time resolutions as low as 150 ps full width at half maximum (FWHM) with LYSO readout have been reported.
In some of the earlier PET systems, gamma detectors could be used only to determine the location of gamma interaction with the detector in two dimensions, which gave rise to parallax errors. More particularly, a conventional two-dimensional measurement of the spatial location of a detected gamma ray absorption event in the scintillating crystal is limited to a two-dimensional point in the X, Y plane of the crystal. The paralax error is a key limiting factor in image resolution, especially for larger patients. This reduces the detectability of small lesions in the outer field of view.
The depth of interaction (DOI) is an important parameter when applied to imaging detector geometries where the directions from which incident gamma rays impinge upon the crystal are not all substantially normal to the crystal surface. If incident gamma rays intersect the crystal from directions not normal to the crystal, the unknown depth of interaction of those gamma rays within the crystal will result in an additional uncertainty in the measured position of the interaction because of the parallax effect, if only a two dimensional (i.e., X, Y) spatial location is calculated for such an absorption event. A detailed explanation of the importance of and the problems associated with the DOI is provided in “Maximum Likelihood Positioning in the Scintillation Camera Using Depth of Interaction,” D. Gagnon et al., IEEE Transactions on Medical Imaging, Vol. 12, No. 1, March 1993, pp. 101-107.
Thus, parallax errors could be reduced by using DOI information to increase the spatial resolution of the system, i.e., to provide the location of gamma interaction in three dimensions in space. In this regard, some PET scanners are able to provide DOI information using axially stacked scintillators which use a pulse shape discrimination method to minimize parallax error. For this example, a DOI detector includes at least two different types of crystal materials, each of which has a different scintillation decay constant, arranged in multiple layers. By discriminating based on the pulse shape, one can differentiate between interaction events that take place in either crystal layer. The layers are subdivided into individual pixel elements to discriminate where within a given layer the interaction has taken place, and reflector partitions may be provided between the crystal elements to better identify the crystal elements in which the interactions take place. Furthermore, a light guide, with or without grooves, and photosensors (e.g. PMTs or other solid state devise) are employed on a single side of the detector in conventional manner.
FIG. 2 shows a diagram of radial spatial resolution vs. radial distance with a simulated resolution across axial field of view in case of a double layer DOI detector (“2 DOI”) and a non-DOI detector (“0 DOI”). So far, conventional PET systems (“0 DOI”) suffer from DOI problems and the resolution deteriorates for larger objects across the axial field of view. Therefore, lesions in the outer contour of the human body (especially large patients) like cancerous lymph nodes cannot be detected with the same image resolution as objects in the body centre. Multi layer DOI detectors (“2 DOI”) could reduce this effect drastically, increase the small lesion detectability and increase life expectancy of the patients.
FIG. 3 shows a schematic example of a double layer DOI detector with shifted pixels in horizontal direction (i.e. x-direction of the coordinate system shown in FIG. 1) for detecting incident gamma photons xx. The double layer DOI detector with single sided readout and depth encoding is composed of two layers of scintillator arrays 22, 24. Each layer is composed of a plurality of polished crystals 10 with a small pitch (e.g. about 1 mm). This assembly may be mounted with a light guide 25 onto a dSiPM 28 array with a small pitch (e.g. about 4 mm). The width of the detector block may be about 32 mm. However, such multi layer DOI encodings suffer from deterioration in timing resolution, as the light is spread over a larger area. This is especially problematic for analog SiPM readout due to their much higher noise floor than photo multiplier tubes (PMTs), which is 1000-10000 times higher per unit area. Therefore, excellent timing can only be achieved when a small sensitive area (at low sensor temperature) is used. The digital SiPMs overcome this problem as individual microcells can be deactivated to reduce the noise floor by several decades.